The present invention relates to a measuring device for measuring radiation of which radiation quantums occur in the Poisson distribution, such as X rays or .gamma. (gamma) rays.
It is a common practice that a dosage of incident radiation such as X rays or .gamma. (gamma) rays is measured by directly or indirectly counting photons of the incident radiation. In a typical indirect photon measuring method, an ionization chamber is used as a radiation detector. The photons are injected into the ionization chamber to ionize ionization gas contained therein. The ionization charges are extracted in the form of current. The amount of the current indicates an incident radiation dosage.
When a frequency of the arrivals of photons is low, the output signal from the radiation detector is detected in the form of a pulse each time that a photon arrives in the radiation detector or the ionization chamber. These pulses can be counted directly. In the case of the radiation with a low frequency of arrival photons, photons are individually detected by the abovementioned method and is directly subjected to a calculation for the radiation measurement.
A tomography apparatus called a computerized tomography (referred to as a CT scanner) has been known as a radiation diagnosis apparatus. The apparatus is provided with an X ray source 1 for radiating a flat fan beam X ray Fx, and a radiation detector 2 with a plurality of radiation detector elements D for detecting the X ray, disposed side by side, as shown in FIGS. 1A and 1B. The X ray source 1 and the detector 2 are disposed opposite one another with an object P under exposure interposing therebetween. In operation, the X ray source 1 and the detector 2 are rotated about the object P in the same direction and at the same angular speed to collect the X ray radiation data on the cross sections of the object as viewed in various directions. The collected data is analyzed by an electronic computer to compute X ray absorptivity at individual positions of each cross section. The graduations are formed corresponding to the absorptivities to reconstruct picture information of the cross sections of the object. This apparatus can provide clear tomographies of the tissues from soft tissue to hard tissue.
The X ray detector 2 includes a number of radiation detecting cells each forming an ionization chamber filled with high pressure gas such as Xe (xenon). The detector detects the energy of X rays transmitted through the object P to produce an ionization current. The current is used as the detected data obtained by the X ray projection to the object.
For collecting the X ray radiation data, the X rays from the X ray source 1 are projected to the object P located between the X ray source 1 and the X ray detector 2. Photons in the radiation from the radiation source 1 oriented at a given angle are transmitted through the object P and enter the detector 2, through paths connecting each detector cell with the X ray source (referred to as an X ray path). In the detector 2, the photons impinge on the high pressure gas in the cells of the detector 2 to ionize the gas. The ionization charges are detected in terms of ionization current. The current is then integrated for a given period of time. The integrated value is discharged through a discharge circuit with a given time constant. The discharge time is used for the X ray radiation data on each X ray path related to each detector cell. After the data collection on every X ray path at the angular position is completed, the data collection on every X ray path at the next angular position starts.
In the CT scanner, the X ray absorptivity of the object depends largely on the arrival photon energy. The energy spectrum of the arrival photons is widely spread. For this reason, the picture image obtained by the method described above is different from that obtained by directly counting the arrival photon for providing the X ray absorption data. A simulation shows that when an object with tissues of which the X ray absorption are approximate to each other are photographed, they have a low contrast. Thus, if the indirect measuring method, in which the ionization charge is extracted and the radiation dosage is detected in the form of the electrical quantity, is used, an image formed is not clear. When the direct measuring method is employed, on the other hand, the image is clearly formed. For this reason, if the direct measuring method can be used, the advantages resulting from it are very great.
In the CT scanner, the frequency of the arrival of photons is very high and a great number of photons must be measured. In this respect, it has been difficult to detect and directly count individual photons, when considering the operating speed of the radiation detector and the counter at the present stage of development. To cope with this technical problem, the prior art treats randomly arriving photons as a continuous stream of photons. Accordingly, the output signal from the detector is produced in the form of a continuous stream of electrons, i.e. a current. In the prior art, the current output signal is amplified by a preamplifier, sampled for a given period of time (several msec.) and is integrated by an integrator. The output signal from the integrator is A/D converted into digital data and then applied to an electronic computer.
A scheme of such a photon detecting system is shown in FIG. 2. The radiation detector 2 is made up of a plurality of detecting cells D1 to Dn each of which is of the ionization chamber type, has a low response, and produces arrival radiation dosage in the form of current. The output currents from the individual cells D1 to Dn are amplified by preamplifiers AMP1 to AMPn, which are then integrated by integrators INT1 to INTn. The output signals from the integrators INT1 to INTn are selectively extracted by a multiplexer of which the output signal is in turn applied to an A/D converter ADC where it is converted into digital data. The digital data are applied as X ray absorption data to a host system such as an electronic computer.
Another prior art devices uses smoothing filters for the integrators. The output signals from the filters are sampled at fixed periods and applied to an A/D converter.
Another prior art device is illustrated in FIG. 3. As shown, the integrator INT1 is of the double integration type. The output signal from the integrator INT1 is compared with a reference voltage Vref by a comparator CMP. The output signal from the comparator CMP is taken out through an AND gate AND1. The output signal from the AND gate AND1 is used as a control signal for an AND gate AND2. With the control signal, the AND gate AND2 is controlled to allow a clock pulse to pass therethrough. The clock pulse is counted by a counter CTR to provide a count corresponding to the detected radiation dosage. The count value is used as digitalized X ray absorption data.
A switch SWa for switching an input signal is provided at the input stage of the integrator INT1. During an X ray radiation period of one pulse, the switch SWa is closed to store the output signal from the detecting element D into an integration capacitor C. The charge in the capacitor C is applied to a comparator CMP where it is compared with a reference voltage Vref. When the former exceeds the latter, the comparator CMP produces an output signal.
The output signal from the comparator is applied, during the X ray radiation rest period, to the AND gate AND1 controlled by a control signal transferred from a control system (not shown). Immediately after the X ray radiation for one minute ends, the output signal from the comparator is produced through the AND gate AND1.
Connected to the integrator INT1 are a DC power source V for feeding a charge with the opposite polarity to the integration capacitor C and a constant current source including a resistor R and a switch SWb. The switch SWb is closed by the output signal from the AND gate AND1. Accordingly, immediately after the X ray radiation ends, the switch SWb is closed and a constant current is fed to the input side of the integrator INT1. Then, the charge stored in the integration capacitor C is discharged according to a discharge characteristic.
As a result, the output level of the integrator INT1 drops to below a reference voltage Vref. Then, the output signal from the comparator CMP disappears and the output signal from the AND gate AND1 also disappears.
During this period, the output signal from the AND gate AND1 is applied as a gate control signal to the AND gate AND2. Accordingly, the AND gate AND2 applies the clock pulses received to the counter CTR during the gate controlled period. As a result, the counter CTR counts a count corresponding to the stored charge in the integrator INT1 to provide digital X ray absorption data corresponding to the arrival X ray dosage.
In the measuring systems shown in FIGS. 2 and 3, data is treated in an analog form in the processing system from the detector to the A/D converter. This fact is accompanied by the following problems.
(1) The photon data of photons randomly occurring, which is essentially digital data, is treated as analog data in most of the data processing system except the final stage of the process. PA1 (2) Unless the noise in the analog system is be suppressed to an extreme (e.g. 10 to 30 uV-rms), when the X ray dosage is low or the attenuation of photons transmitted through the object is high, the noise in the analog system is relatively distinctive, and may damage the CT image quality. It is desirable that the S/N ratio of the CT image achieves a physical limitation determined by the number of photons. PA1 (3) Unless the photon-charge conversion coefficient and a signal transmission function in the signal processing system from the detector to the signal processing system is reduced to a minimum, a measurement error occurs, so that an artifact deteriorates the sharpness of the image or the reliability of the measured values is reduced. PA1 (4) Generally, even when the number of photons striking the detector is zero, the output signal from the processing system (the output signal from the A/D converter) is not zero because of the presence of the dark current and the off-set voltage of the amplifier. In order to prevent the error by the dark current or the off-set voltage, the output signal from the processing system, when its input is zero. is repeatedly measured at a short cycle to measure the offset data, and the off-set data is subtracted from the output signal of the processing system when the photon input is received. PA1 (5) The arrival photons are not monocolor but multicolor and hence have widely distributed energy spectrums. As for individual photons, the shorter the wavelength of the photon, the larger the output charge or output current from the detector becomes. As for the difference in the absorption coefficients of the tissues in the object, it becomes more intensified as the energy of the photon goes lower. For this reason, in simple X-ray photographing, it is desirable that the X-rays at low energy (longer wavelength) should be used for photographing the soft tissue. However, in a conventional CT, the output signal from the measuring system is greatly influenced by the photons at high energy, while it is less influenced by the photons at low energy containing precise information of a minute difference in X ray absorptions of the tissues in the object. As a result, a resolution of the CT image at a low contrast is reduced.
These problems can be solved if the arrival photons can be individually detected and counted.
To detect and count individual photons of the X ray at a high dose rate, the operating speed of the detector and the counter must be extremely high. The arrival rates of photons arriving at each detector of some typical X ray CT scanners are given in Table 1.
TABLE 1 ______________________________________ Number of Objects photons/sec ______________________________________ A None (All air) 4 .times. 10.sup.8 B Test phantom with a container of 4 .times. 10.sup.7 120 mm in diameter containing water C Test phantom with a container of 4 .times. 10.sup.6 240 mm in diameter containing water ______________________________________
A continuous X ray at 120 KV/20 mA was used for an X ray source when creating the above table. When an organism is used for the object, the measuring error must be within about .+-.1%. This tolerance is selected so as not to disturb a physical limitation of the S/N ratio determined by the number of the arrival photons. The number of arrival photons in the case of the head of a human is substantially equal to that in the case of the water contained in the 120 mm container, and is about 4.times.10.sup.7 times/sec. For detecting all the arrival photons, the detector and counter are required to operate following the occurence at the photons of 4.times.10.sup.7 times/sec. Actually, a distribution of the photons is random and not uniform. Therefore, the measuring system operating simply at 40 MHz fails to detect all of the respective photons.
This phenomenon will be described in more detail referring to FIGS. 4A and 4B, 5A to 5C, FIG. 6 and FIG. 7. It is assumed that photons arrive at the detector one for t seconds and at random periods, as shown in FIG. 4A. The response of the detector is relatively slow to such an extent that it takes at least t seconds to count the succeeding a second photon. Accordingly, photon arriving within the interval t is treated as if it is a part of the preceeding photon, and is made to correspond to one pulse. This phenomenon is called a "bunching" and frequently occurs in the detector of the CT scanner. In this explanation, the detector contains a discriminator for discriminating the arrival photons from the noise. As seen from FIG. 4B, showing the output signal from the detector, the detector can detect only seven photons of the 14 arrival photons, even if the counting speed is high.
FIGS. 5A to 5C illustrate waveforms in a case where the response speed of the detector is ideal but the counter's operating speed is low. FIG. 5A illustrates a distribution of arrival photons, and FIG. 5B shows a waveform of the input signal to the counter. For executing the incremental operation of the counter, the high level period of the signal for each count is t.sub.H or more and the low level period is t.sub.L or more. The counter must see a pulse duration of t.sub.H or greater followed by a delay of t.sub.L or greater in order to count one photon. Therefore, when the pulse is a positive pulse, as shown in FIG. 5A, the positive level period must be continued for the time period t.sub.H or more, as shown in FIG. 5B. If the pulse is shaped such that the positive level period of the output signal from the detector continues over the time t.sub.H, the counter counts only seven photons, which is the half of the total arrival photons, for one cycle of counting operation, as described referring to FIGS. 4A and 4B.
In any of these cases shown in FIGS. 4A and 4B, and FIGS. 5A to 5C, the state of bunching varies due to the instability of the response of the detector or the set time t.sub.H drifts. As a result, a characteristic of the detecting and counting system is that a % of the input photons are dropped in counting transitions, so that the measuring sensitivity drifts even though the measuring system is of the digital type.
FIG. 6 shows the value of a counting coefficient K (i.e. a ratio of the number of the arrival photons to the count of the photons by the counter) when a number M of events, e.g. the average number of photons arriving during a minimum operation cycle (10 ns) of a prior art device. In the case of FIG. 5, the characteristic curve is expressed by K=e.sup.-M where e is a natural logarithm.
FIG. 7 shows count values R when a circuit with a minimum operation cycle of 10 ns counts a number M of random events, which occur during 1 ms. As seen from FIGS. 6 and 7, in the prior art device, as the average rate of the photons arriving during the minimum operation cycle increases, the detecting and counting systems fail to follow the occurence of the photons, so that the number of the count is abruptly decreased resulting in incorrect counting.